Ultrasound is widely used in modern medicine for diagnostics and minimally invasive treatment in such fields as obstetrics, cardiology, endocrinology, gastroenterology, neurology, ophthalmology, urology, osteoporosis, and clinical diagnostics. Ultrasound diagnostics uses low-power ultrasonic scanners for investigation and visualization of inner organs, tissue layers and structures, for determination of blood flow direction and velocity, for measurement of density and other parameters of tissues, and for detection of cancer and other tumors. In diagnostics, acoustic lenses have been traditionally used in pulse mode to manipulate the wave front propagation delays. In therapeutic applications, continuous ultrasound waves with an average acoustic intensity of up to several watts per centimeter square at the transducer surface are typically used to focus ultrasound. The focused ultrasonic waves produce highly localized and intense acoustic fields, up to several hundreds of watts in power density, and enable controlled, deep-reaching and localized treatment of malignant tissues, with few secondary effects for surrounding health tissues. It is beneficial to control ultrasonic energy deposition for quickly overheating target focal tissue while minimizing the impact on surrounding non-targeted tissues. The mastery of focusing determines the success of therapy and requires an understanding of the vibration condition of the radiating surface and thermal and mechanical constraints. Because acoustic focusing is an interference phenomenon, the phase of individual ultrasound rays becomes a controlling factor in a continuous therapy mode. In a diagnostic imaging mode, focusing limits the beam width and constrains the acoustic energy content of the beam to a smaller cross sectional area, hence improving imaging sensitivity. In this mode, the beam is typically focused using a fixed lens that just bends acoustic rays and preserves the pressure-time waveform of incoming signals. Imaging lenses are used in pulsed mode where their function relies primarily on determining and manipulating the wave front propagation delays. For therapeutics, the mode of operation is typically continuous wave, in which case the phase becomes an important lens design factor as opposed to wave front propagation delay. Traditional convex or concave lenses (Folds, Focusing properties of solid ultrasound cylindrical lenses, 53, 3, pp 826-834, 1973) that converge light rays towards the lens principal axis offer a simple method to focus low power acoustic energy in both therapy and imaging. However, high acoustic absorption in thicker regions of the lenses and excessive heat build up result in a poor lens longevity and large focusing aberration when attempts are made to focus high power acoustic energy in a continuous regime. Hence, thin focusing lenses with discrete phase shifts are both permissible and beneficiary in therapy, greatly reducing overall lens depth profile and allowing different designs, including zone plate Fresnel (Hadimioglu et al, 1993), multilevel (Chan et al, Finite element analysis of multilevel acoustic Fresnel lenses, Vol 43, 4, 1996), field conjugate (Lalonde and Hunt, Variable frequency field conjugate lenses for ultrasound hyperthermia, 42, 5, 825-831, 1995) and other designs (Rosenberg, High intensity ultrasound, Moscow, pp 69-91, 1949; Tarnoczy, Sound focusing lenses and waveguides, Ultrasonics, 115-127, 1965).
Discrete phase acoustic focusing lenses in combination with flat transducers or arrays offer an elegant and cost effective solution for hyperthermia treatment of cancer and tumors, where the tissue is heated using ultrasound to temperatures of 43-45° C. for several minutes. It is well known that tumor cells become much more susceptible to radiotherapy and chemotherapy under elevated temperature. In physiotherapy lens focused ultrasound may be used to increase the elasticity of sinews and scars, improve the mobility of joints, provide analgesic effects, alter blood flow, and produce muscular spasms. High intensity ultrasound (10-2000 W/cm2) is used for tissue ablation, cutting, fractionation (histotripsy) and for arresting internal bleeding (hemostasis). Historically, piezoelectric and magnetostrictive transducers are widely used to transform generate a high intensity ultrasound field.
In therapeutic applications the precision targeting of deep tissues is important. Desired therapeutic effect must be confined to a small spot within the body where temperature elevation is sufficient to create a localized tissue impact without affecting surrounding tissue and organs. This technique is used to selectively destroy the unwanted tissue within the body without perturbing adjacent tissues. Typically, heating the tissue to 60° C.-80° C. results in tissue necrosis, a process commonly termed as thermal ablation. In most cases, the high intensity focused ultrasound is used in thermal ablation procedures. Ultrasound focusing can be achieved by having concave focused transducers producing convergent beams of predetermined geometry and/or by manipulating the driving electrical signals (phase and amplitude) of multiple active transducers (Cathignol, 2002, High Intensity Piezoelectric Sources for Medical Applications: Technical Aspects, Nonlinear Acoustics at the Beginning of the 21st Century, 1, 371-378.). Single focused elements are more economical but require mechanical steering and suffer a loss of acoustic efficiency due to heating and presence of parasitic surface waves (Kluiwstra et al., 1997, Design Strategies for Therapeutic Ultrasound Phased Arrays, SPIE International Medical Imaging Symposium, Chapelon et al. Transducers for therapeutic ultrasound, Ultrasound in Med. & Biol., Vol. 26, No. 1, pp. 153-159, 2000).
Ultrasound systems use relatively small, low-power transducers for diagnostic visualization and large high-power transducers for therapy. Typically, the radiation surfaces of the two types of transducers coincide and often form a surface of revolution of a conic section: circle, ellipse or parabola. Transducers with large radiating surfaces are used to generate sufficient acoustic power and are expensive to manufacture. Additionally, the applicability of large concave transducers is limited to an open field clinical cases, where the size of the transducer does not matter, as opposed to the most intra-luminal or intra-cavity applications, where access is limited and the dimensional requirements counter acoustic power and sensitivity requirements.